Ultrasound imaging system

ABSTRACT

The present invention is directed to an ultrasound imaging system and method for Doppler processing of data. The ultrasonic imaging system efficiently addresses the data computational and processing needs of Doppler processing. In a preferred embodiment, the ultrasound imaging system of the present invention includes a processing module; and memory operable coupled to the processing module, wherein the memory stores operational instructions that cause the processing module to map serial data to vector representation, calculate an auto-correlation function of the data, calculate a phase shift of the auto-correlation function to generate a monotonic function covering all values of the phase shift corresponding to a range of Doppler velocities and display the resultant images, for example, as color images.

CROSS REFERENCE TO RELATED APPLICATIONS

The present application is a continuation of U.S. application Ser. No.11/807,236 filed May 25, 2007, which claims priority to U.S. applicationSer. No. 10/259,251 filed Sep. 27, 2002, now U.S. Pat. No. 7,223,242 andU.S. application Ser. No. 09/966,810 filed Sep. 28, 2001, now U.S. Pat.No. 6,638,225. The entire contents of the above patents and applicationsare incorporated herein by reference.

BACKGROUND OF THE INVENTION

Ultrasonic diagnostic equipment has become an indispensable tool forclinical use. For approximately the past twenty years, real-time B-modeultrasound imagers are used for investigating all soft tissue structuresin the human body. One of the recent developments within medical imagingtechnology is the development of Doppler ultrasound scanners.

Doppler ultrasound is an important technique for non-invasivelydetecting and measuring the velocity of moving structures, andparticularly to display an estimate of blood velocity in the body inreal time.

The basis of Doppler ultrasonography is the fact that reflected and/orscattered ultrasonic waves from a moving interface undergoes a frequencyshift. In general the magnitude and the direction of this shift providesinformation regarding the motion of this interface. How much thefrequency is changed depends upon how fast the object or movinginterface is moving. Doppler ultrasound has been used mostly to measurethe rate of blood flow through the heart and major arteries.

There are several forms of depiction of blood flow in medical Dopplerimaging or more generally different velocity estimation systems thatcurrently exist: Color Flow imaging, power Doppler and Spectralsonogram. Color flow imaging (CFI), interrogates a whole region of thebody, and displays a real-time image of mean velocity distribution. CFIprovides an estimate of the mean velocity of flow with a vessel by colorcoding the information and displaying it, super positioned on a dynamicB-mode image or black and white image of anatomic structure. In order todifferentiate flow direction, different colors are used to indicatevelocity toward and away from the transducer.

While color flow imaging displays the mean or standard deviation of thevelocity of reflectors, such as the blood cells in a given region, powerDoppler (PD) displays a measurement of the amount of moving reflectorsin the area, similarly to the B-mode image's display of the total amountof reflectors. A power Doppler image is an energy image in which theenergy of the flow signal is displayed. Thus, power Doppler depicts theamplitude or power of the Doppler signals rather than the frequencyshift. This allows detection of a larger range of Doppler shifts andthus better visualization of small vessels. These images give novelocity information, but only show the direction of flow. In contrast,spectral Doppler or spectral sonogram utilizes a pulsed wave system tointerrogate a single range gate or sampling volume, and displays thevelocity distribution as a function of time. The sonogram can becombined with the B-mode image to yield a duplex image. Typically, thetop side displays a B-mode image of the region under investigation, andthe bottom displays the sonogram. Similarly, the sonogram can also becombined with the CFI or PD image to yield a triplex image. The time fordata acquisition is then divided between acquiring all three sets ofdata, and the frame rate of the images is typically decreased, comparedto either CFI or duplex imaging.

The current ultrasound systems require extensive complex data processingcircuitry in order to perform the imaging functions described herein.Doppler processing for providing two-dimensional depth and Dopplerinformation in color flow images, power Doppler images and/or spectralsonograms require millions of operations per second. There exists a needfor an ultrasound imaging system that provides for compute-intensivesystems and methods to efficiently address the data processing needs ofinformation, such as Doppler processing.

SUMMARY OF THE INVENTION

The present invention is directed to an ultrasound imaging system andmethod for Doppler processing of data. The ultrasonic imaging systemefficiently addresses the data computational and processing needs ofDoppler processing. Software executable sequences in accordance with apreferred embodiment of the present invention determines the phase shiftand the auto-correlation phase of filtered image data. In a preferredembodiment, the system of ultrasonic imaging also includes a sequence ofinstructions for Doppler processing that provides the functions fordemodulation, Gaussian Match filtering, auto-correlation calculation,phase shift calculation, frame averaging, and scan conversion.

In a preferred embodiment, the processing system includes parallelprocessing elements which execute Single Instruction Multiple Data(SIMD) or Multiple Instruction Multiple Data (MIMD) instructions. Acomputer having a Pentium® III processor including MMX™ technology is anexemplary computational device of a preferred embodiment of theultrasonic imaging system in accordance with the present invention.

A method of the present invention includes imaging a region of interestwith ultrasound energy using a portable ultrasound imaging system whichin turn includes a transducer array within a handheld probe. Aninterface unit is connected to the handheld probe with a cableinterface. The interface unit has a beamforming device connected to adata processing system with another cable interface. Output signals fromthe interface unit are provided to the handheld probe to actuate thetransducer array, which in turn delivers ultrasound energy to the regionof interest. The ultrasound energy returning to the transducer array iscollected from the region of interest and transmitted from the handheldprobe to the interface unit. A beamforming operation is performed withthe beamforming device in the interface unit. The method furtherincludes transmitting data from the interface unit to the dataprocessing system such that the data processing system receives abeamformed electronic representation of the region of interest. The dataprocessing system has at least one parallel processing elementintegrated with a microprocessor to execute a sequence of instructionsfor Doppler processing and displaying of Doppler images.

In a preferred embodiment, the ultrasound imaging system of the presentinvention includes a processing module; and memory operable coupled tothe processing module, wherein the memory stores operationalinstructions that cause the processing module to map serial data to avector representation, demodulate the data to obtain in-phase andquadrature sample data, calculate an auto-correlation function of thedata, calculate a phase shift of the auto-correlation functionrepresented as a monotonic function in the interval corresponding to therange of Doppler velocities according to the Nyquist criterion andexpressed as a simple mathematical function, convert the phase shift toan index and display the images, for example, as color images.

Preferred embodiments of the present invention include the portableultrasound system comprising one of at least a color Doppler mode, adirectional power Doppler mode, a power Doppler mode and a pulsed waveDoppler mode. The portable ultrasound system includes adjustablecontrols for one of at least scan area selection, velocity display,steering angles, color inversion, color gain, color priority, colorpersistence, color baseline and frame rate.

The foregoing and other features and advantages of the system and methodfor ultrasound imaging will be apparent from the following moreparticular description of preferred embodiments of the system and methodas illustrated in the accompanying drawings in which like referencecharacters refer to the same parts throughout the different views.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a block diagram of a preferred embodiment of the ultrasoundimaging system in accordance with the present invention;

FIGS. 2A and 2B are flow charts of a preferred embodiment of a methodfor Doppler processing in accordance with the present invention;

FIG. 3 is a diagram illustrating a preferred embodiment of a method ofdata mapping for parallel computation in accordance with the presentinvention; and

FIGS. 4A and 4B are graphical representations of the Doppler phase shiftcalculations in accordance with preferred embodiments of the presentinvention.

FIG. 5 is a view of a display screen showing an ultrasound imageillustrating the color Doppler mode in accordance with a preferredembodiment of the present invention.

FIGS. 6A and 6B illustrate a view of a display screen showing the colorDoppler mode of an ultrasound image and the available controls,respectively, in accordance with a preferred embodiment of the presentinvention.

FIGS. 7A-7D illustrate the controls and image sizes available byadjusting the scan area, respectively, in accordance with a preferredembodiment of the present invention.

FIGS. 8A-8D illustrate the controls available for adjusting differentparameters such as pulse repetition frequency, wall filter, steeringangle and color inversion for a color Doppler mode in accordance with apreferred embodiment of the present invention.

FIGS. 9A and 9B illustrate the color Doppler reference bar that is usedfor a color invert adjustment in accordance with a preferred embodimentof the present invention.

FIG. 10 illustrates additional adjustments available to a user in thecolor Doppler mode such as color gain, color priority, color persistenceand color baseline in accordance with a preferred embodiment of thepresent invention.

FIGS. 11A and 11B illustrate the images and controls provided bydirection power Doppler, respectively, in accordance with a preferredembodiment of the present invention.

FIGS. 12A and 12B illustrate an image and the controls provided to auser in a power Doppler mode in accordance with a preferred embodimentof the present invention.

FIGS. 13A-13C illustrate a real-time mixed mode (B-mode scan), a pulsedDoppler waveform and the controls provided to a user in a pulsed waveDoppler mode in accordance with a preferred embodiment of the presentinvention.

The drawings are not necessarily to scale, emphasis instead being placedupon illustrating the principles of the invention.

DETAILED DESCRIPTION OF THE INVENTION

The ultrasound imaging system is directed at a Doppler processing systemin a portable ultrasound system. In a preferred embodiment, theultrasonic imaging system includes parallel computation units and amemory having stored therein instructions to process data and displayultrasound images using computer-efficient methods.

A preferred embodiment of the ultrasound imaging system includes apulse-Doppler processor for color flow imaging or map applications.Color flow (CF) imaging combines in a single modality the abilities ofultrasound to image tissue and to investigate blood flow. CF imagesconsist of Doppler information that can be color-encoded andsuperimposed on a B-mode gray-scale image.

Color-flow imaging is a mean velocity estimator. There are two knowndifferent techniques for computing the mean velocity. First, in a pulsedDoppler system, Fast-Fourier Transforms (FFTs) can be used to yield thevelocity distribution of the region of interest, and both the mean andthe variance of the velocity profile can be calculated and displaced asa color flow imaging. The other approach is the one dimensionalauto-correlation technique described by Kasai et al in “Real-Time TwoDimensional Blood Flow Imaging Using an Auto-correlation Technique” inthe IEEE Transactions on Sonics and Ultrasonics Vol. SU-32, No. 3 in May1985, the entire contents of which are incorporated herein by reference.

Mean blood flow velocity is estimated from the frequency spectra ofechoes. An estimate of the mean velocity in the range gate or samplevolume gives an indication of the volume flow rate. As the frequency ofthe range or depth gated and sampled signal is proportional to thevelocity, the spatial mean velocity can be determined by the meanangular frequency of P(ω) and is expressed as:

$\begin{matrix}{\varpi = \frac{\int_{- \infty}^{+ \infty}{\omega \; {P(\omega)}\ {\omega}}}{\int_{- \infty}^{+ \infty}{{P(\omega)}\ {\omega}}}} & (1)\end{matrix}$

where P(ω) is the power density spectrum of the received, demodulatedsignal. Equation (1) gives the mean Doppler frequency shift due to theblood flow. The mean blood flow velocity v can then be estimated by thefollowing equation:

$\begin{matrix}{\overset{\_}{v} = {\frac{\varpi}{\omega_{0}}\frac{c}{2\mspace{11mu} \cos \mspace{11mu} \theta}}} & (2)\end{matrix}$

where c is the velocity of sound and θ the angle between the sound beamand the blood flow vector.

The extent of turbulence in blood flow may be inferred from the varianceof the spectrum. Since the Doppler frequency directly relates to theflow vector, i.e., flow direction and speed, in an ultrasonic samplevolume, the spectrum spread broadens in accordance with flowdisturbance. While in laminar flow, the spectrum spread is narrow, sincea uniform flow vector gives a singular Doppler frequency shift. The meanangular frequency can be determined by the phase-shift ofauto-correlation of the complex signal z(t). The inverse Fouriertransform of the power density spectrum is the auto-correlation functionR(τ) and is expressed as:

$\begin{matrix}{{R(\tau)} = {{\int_{- \infty}^{+ \infty}{{P(\; \omega)}^{j\omega\tau}\ {\omega}}} \equiv {{A(\tau)}^{{j\varphi}{(\tau)}}}}} & (3)\end{matrix}$

From the moment's theorem of Fourier transforms, it can be shown that

$\begin{matrix}{{{\overset{.}{R}(0)} = {j{\int_{- \infty}^{+ \infty}{\omega \; {P(\omega)}{\omega}}}}}{and}} & (4) \\{{R(0)} = {\int_{- \infty}^{+ \infty}{{P(\omega)}{\omega}}}} & (5)\end{matrix}$

It follows then

$\begin{matrix}{\overset{\_}{\varpi} = \frac{\overset{.}{R}(0)}{j\; {R(0)}}} & (6)\end{matrix}$

Therefore, the mean velocity estimation can be reduced to an estimationof the auto-correlation and the derivative of the auto-correlation. Theestimator given by the above expression can be calculated when data fromtwo returned lines are used. From Equation (3),

{dot over (R)}(0)=jA(0){dot over (φ)}(0) and R(0)=A(0)  (7)

Substituting the above equations into (6), we have

$\begin{matrix}{\overset{\_}{\varpi} = {{{\overset{.}{\varphi}(0)} \approx \frac{{\varphi (1)} - {\varphi \left( {- 1} \right)}}{2}} = {\varphi (1)}}} & (8)\end{matrix}$

Generally, φ(1) can be determined by either of the following methods

$\begin{matrix}{{\varphi (1)} = {\arctan \left( \frac{{Im}\left\{ {R(1)} \right\}}{{Re}\left\{ {R(1)} \right\}} \right)}} & (9) \\{{\varphi (1)} = {\arcsin \left( \frac{{Im}\left\{ {R(1)} \right\}}{{R(1)}} \right)}} & (10) \\{{\varphi (1)} = {\arccos \left( \frac{{Re}\left\{ {R(1)} \right\}}{{R(1)}} \right)}} & (11)\end{matrix}$

In a preferred embodiment of the ultrasonic imaging system,

${{sign}\left( {\varphi (1)} \right)}{\sin^{2}\left( \frac{\varphi (1)}{2} \right)}$

represents the phase-shift, where

$\begin{matrix}{{{sign}(x)} = \left\{ {\begin{matrix}1 & {x > 0} \\0 & {x = 0} \\{- 1} & {x < 0}\end{matrix}{{sign}\left( {\varphi (1)} \right)}{\sin^{2}\left( \frac{\varphi (1)}{2} \right)}} \right.} & (12)\end{matrix}$

is a monotonic function of φ(1) in the interval (−π, +π). Thus, everyvalue of

${sign}\left( {\varphi (1)} \right){\sin^{2}\left( \frac{\varphi (1)}{2} \right)}$

uniquely defines a φ(1) in the interval (−π, +π) and vice versa.

Further,

$\begin{matrix}{{{sign}\left( {\varphi (1)} \right)} = {{{sign}\left( {\sin \left( {\varphi (1)} \right)} \right)} = {{sign}\left( {{Im}\left\{ {R(1)} \right\}} \right)}}} & (13) \\{{\sin^{2}\left( \frac{\varphi (1)}{2} \right)} = {\frac{1 - {\cos \left( {\varphi (1)} \right)}}{2} = \frac{1 - \frac{{Re}\left\{ {R(1)} \right\}}{{R(1)}}}{2}}} & (14)\end{matrix}$

And therefore:

$\begin{matrix}{{{s{ign}}\left( {\varphi (1)} \right)} = {{\sin^{2}\left( \frac{\varphi (1)}{2} \right)} = {\frac{1}{2}{{{sign}\left( {{Im}\left\{ {R(1)} \right\}} \right)}\left\lbrack {1 - \frac{{Re}\left\{ {R(1)} \right\}}{{R(1)}}} \right\rbrack}}}} & (15)\end{matrix}$

Similarly, the φ(1) can be determined based on

$\begin{matrix}{{\tan \left( \frac{\varphi (1)}{2} \right)} = {\frac{\sin \mspace{11mu} {\varphi (1)}}{1 + {\cos \left( {\varphi (1)} \right)}} = \frac{\frac{{Im}\left\{ {R(1)} \right\}}{{R(1)}}}{1 + \frac{{Re}\left\{ {R(1)} \right\}}{{R(1)}}}}} & \left( {15a} \right)\end{matrix}$

In a preferred embodiment of the present invention, the estimator givenby the above expressions 15 and/or 15(a) can be calculated when datafrom at least two returned lines are used. The equations 15 and 15arepresent the phase shift φ as a monotonic function in the intervalbetween negative and positive pi (−π, +π) and express the phase shift assimple mathematical calculations. In a preferred embodiment of thepresent invention, more lines are used in order to improve thesignal-to-noise ratio. Data from several RF lines are needed in order toget useful velocity estimates by the auto-correlation estimator.Preferably, in a particular embodiment between eight (8) and sixteen(16) lines are acquired for the same image direction. The lines aredivided into range gates throughout the image depths, and the velocityis estimated along the lines.

The CFI pulses are interspersed between the B-mode image pulses induplex imaging. It is known that a longer duration pulse train gives anestimator with a lower variation. However, a good spatial resolutionnecessitates a short pulse. In a particular embodiment of the ultrasoundimaging system of the present invention, a separate pulse is preferablyused for the B-mode image, because the CFI pulse is too long for highquality gray-scale image.

While Color Flow Imaging (CFI) sonograph has been an effectivediagnostic tool in clinical cardio-vascular application, power Doppler(PD) imaging provides an alternative method of displaying the bloodstream in the insonified regions of interest. While CF imaging displaysthe mean or standard deviation of the velocity of reflectors such as,for example, blood cells in a given region, PD displays a measurement ofthe amount of moving reflectors in the area, similarly to the B-modeimage's display of the total amount of reflectors. Thus, power Doppleris akin to a B-mode image with the stationary reflectors suppressed.This is particularly useful for viewing small moving particles withsmall scattering cross-sections such as red blood cells.

The power Doppler image displays the integrated Doppler power instead ofthe mean frequency shift as used for color Doppler imaging. As discussedhereinbefore, the color flow mapping is a mean frequency estimation, andcan be expressed as:

$\begin{matrix}{\varpi = \frac{\int_{- \infty}^{+ \infty}{\omega \; {P(\omega)}\ {\omega}}}{\int_{- \infty}^{+ \infty}{{P(\omega)}\ {\omega}}}} & (1)\end{matrix}$

where ω represents the mean frequency shift and P(ω) is the powerdensity spectrum of the received signal. It has also been shownhereinbefore by inverse Fourier transform that

$\begin{matrix}{{R(\tau)} = {\int_{- \infty}^{+ \infty}{P\; (\omega)^{j\omega\tau}\ {\omega}}}} & (16)\end{matrix}$

The total integrated Doppler power can be expressed as

$\begin{matrix}{{pw} = {\int_{- \infty}^{+ \infty}{P\; (\omega)\ {\omega}}}} & (17)\end{matrix}$

Substituting Equation (16) into (17), it follows that

$\begin{matrix}{{R(0)} = {{\int_{- \infty}^{+ \infty}{P\; (\omega){\exp \left( {j\; \omega \; 0} \right)}\ {\omega}}} = {{\int{{P(\omega)}{\omega}}} = {p\; w}}}} & (18)\end{matrix}$

and that the 0th lag of the auto-correlation function can be used tocompute the integrated total Doppler power. In other words theintegrated power in the frequency domain is the same as the integratedpower in the time domain and hence the power Doppler can be computed inone preferred embodiment from either time-domain or the frequency-domaindata. In either embodiment, the unwanted signal from the surroundingtissue such as the vessel walls is removed via filtering.

In a preferred embodiment, the PD computation can be carried out insoftware executing a sequence of instructions on the host processorsimilarly to the computation of the CFI processing describedhereinbefore. In a preferred embodiment, parallel computation units suchas those in the Intel® Pentium® and Pentium® III's MMX™ coprocessorsthat allow rapid computation of the required functions are used.Intel®'s MMX™ is a Pentium® microprocessor that executes applicationsfaster than non-MMX™ Pentium® microprocessor. It is designed to improvethe performance of multimedia and communication algorithms. Thetechnology includes new instructions and data types which achieve higherlevels of performance for these algorithms or host processors. Inparticular, MMX™ Pentium® microprocessors have microprocessorinstructions that are designed to handle video, audio and graphical datamore efficiently. Further, MMX™ technology consists of a SingleInstruction Multiple Data (SIMD) process which makes it possible for oneinstruction to perform the same operation on multiple data items. Inaddition, the memory cache on the MMX™ Processor has increased to, forexample, 32 thousand bytes, which provides for fewer accesses to memorythat is off the microprocessor. In an alternate preferred embodiment, aDigital Signal Processor (DSP) can also be used to perform theprocessing function. Such architecture permits flexibility in changingdigital signal processing algorithms and transmitting signals to achievethe best performance as the region of interest is changed.

As discussed hereinbefore, the frequency content of the Doppler signalcorresponds to the velocity distribution of the blood. It is common todevice a system for estimating blood movement at a fixed depth intissue. A transmitter emits an ultrasound pulse that propagates into andinteracts with tissue and blood. The backscattered signal is received bythe same transducer and amplified. For a multiple-pulsed system, onesample is acquired for each pulse emitted. A display of the distributionof velocities can be made by Fourier transforming the received signaland displaying the result. This display is also called a sonogram. Oftena B-mode image is presented along with the sonogram in a duplex system,and the area of investigation or range gate is displayed on the image.The placement and size of the range gate are determined by the user, andthis determines the time instance for the sampling operation. The rangegate length determines the area of investigation and sets the length ofthe emitted pulse.

The calculated spectral density is displayed on a screen with frequencyon the y-axis and time on the x-axis. The intensity at a point on thescreen indicates the amplitude of the spectrum and is, thus,proportional to the number of blood scatterer moving at a particularvelocity.

FIG. 1 is a schematic functional block of one embodiment of theultrasound imaging system 10 of the invention. Similar imaging systemsare described in U.S. Pat. No. 5,957,846 to Alice M. Chiang et alt,issued Sep. 28, 1999, entitled “Portable Ultrasound Imaging System,” theentire contents of which are being incorporated herein by reference. Asshown, the system 10 includes an ultrasonic transducer array 14 whichtransmits ultrasonic signals into a region of interest or image target12, such as a region of human tissue, and receives reflected ultrasonicsignals returning from the image target. The system 10 also includes afront-end interface or processing unit 18 which is connected by cables16, for example, coaxial cables to the transducer array 14 and includesa transducer transmit/receive control chip 22.

Ultrasonic echoes reflected by the image target 12 are detected by theultrasonic transducers in the array 14. Each transducer converts thereceived ultrasonic signal into a representative electrical signal whichis forwarded to an integrated chip having preamplification circuits andtime-varying gain control (TGC) circuitry 30. The preamplificationcircuitry sets the level of the electrical signals from the transducerarray 14 at a level suitable for subsequent processing, and the TGCcircuitry is used to compensate for attenuation of the sound pulse as itpenetrates through human tissue and also drives the beamforming circuits32 to produce a line image. The conditioned electrical signals areforwarded to the beamforming circuitry 32 which introduces appropriatedifferential delay into each of the received signals to dynamicallyfocus the signals such that an accurate image can be created. Furtherdetails of the beamforming circuitry 32 and the delay circuits used tointroduce differential delay into received signals and the pulsesgenerated by a pulse synchronizer are described in U.S. Pat. No.6,111,816 to Alice M. Chiang et al., issued Aug. 29, 2000 entitled“Multi-Dimensional Beamforming Device,” the entire content of which arebeing incorporated herein by reference.

A memory 30 stores data from a controller 28. The memory 30 providesstored data to the transmit/receive chip 22, the TGC 30 and thebeamformer 32. The output from the system controller 28 is connecteddirectly to a custom or Fire Wire Chipset. The FireWire Chip set isdescribed in co-pending U.S. application Ser. No. 09/966,810 filed Sep.28, 2001, now U.S. Pat. No. 6,638,225, entitled “Ultrasound Probe withIntegrated Electronics,” by Jeffrey M. Gilbert et al., the entirecontents of which are being incorporated herein by reference. “FireWire”refers to IEEE standard 1394, which provides high-speed datatransmission over a serial link. There also exists a wireless version ofthe FireWire standard allowing communication via an optical link foruntethered operation.

The FireWire standard and an ultrasound probe with integratedelectronics as described in co-pending U.S. patent application Ser. No.09/791,491, entitled “Ultrasound Probe With Integrated Electronics,” byAlice M. Chiang et al., the entire contents of which are beingincorporated herein by reference, may be used in preferred embodimentsof the present invention. The FireWire standard is used for multimediaequipment and allows 100-200 Mbps and preferably in the range of 400-800Mbps operation over an inexpensive 6 wire cable. Power is also providedon two of the six wires so that the FireWire cable is the only necessaryelectrical connection to the probe head. A power source such as abattery or IEEE1394 hub can be used. The FireWire protocol provides bothisochronous communication for transferring high-rate, low latency videodata as well as asynchronous, reliable communication that can be usedfor configuration and control of the peripherals as well as obtainingstatus information from them. Several chipsets are available tointerface custom systems to the FireWire bus. Additionally,PCI-to-FireWire chipsets and boards are currently available to completethe other end of the head-to-host connection. CardBus-to-FireWire boardscan also be used.

Although the VRAM controller directly controls the ultrasound scan head,higher level control, initialization, and data processing and displaycomes from a general purpose host such as a desktop PC, laptop, orpalmtop computer. The display can include a touchscreen capability. Thehost writes the VRAM data via the VRAM Controller. This is performedboth at initialization as well as whenever any parameters change (suchas number or positions of zones, or types of scan head) requiring adifferent scanning pattern. During routine operation when data is justbeing continually read from the scan head with the same scanningparameters, the host need not write to the VRAM. Because the VRAMcontroller also tracks where in the scan pattern it is, it can performthe packetization to mark frame boundaries in the data that goes back tothe host. The control of additional functions such as power-down modesand querying of buttons or dial on the head can also be performed viathe FireWire connection.

Although FireWire chipsets manage electrical and low-level protocolinterface to the FireWire interface, the system controller has to managethe interface to the FireWire chipset as well as handling higher levelFireWire protocol issues such as decoding asynchronous packets andkeeping frames from spanning isochronous packet boundaries.

Asynchronous data transfer occurs at anytime and is asynchronous withrespect to the image data. Asynchronous data transfers take the form ofa write or read request from one node to another. The writes and thereads are to a specific range of locations in the target node's addressspace. The address space can be 48 bits. The individual asynchronouspacket lengths are limited to 1024 bytes for 200 Mbps operation. Bothreads and writes are supported by the system controller. Asynchronouswrites are used to allow the host to modify the VRAM data as well as acontrol word in the controller which can alter the operation mode.Asynchronous reads are used to query a configuration ROM (in the systemcontroller FPGA) and can also be used to query external registers or I/Osuch as a “pause” button. The configuration ROMs contain a querible“unique ID” which can be used to differentiate the probe heads as wellas allow node-lockings of certain software features based on a key.

Using isochronous transfers, a node reserves a specified amount ofbandwidth and it gets guaranteed low-overhead bursts of link accessevery 1/8000 second. All image data from the head to the host is sentvia isochronous packets. The FireWire protocol allows for somepacket-level synchronization and additional synchronization is builtinto the system controller.

The front-end processing or interface unit system controller 28interfaces with a host computer 20, such as a desktop PC, laptop orpalmtop, via the custom or FireWire. Chipsets 24, 34. This interfaceallows the host to write control data into the memory 26 and receivedata back. This may be performed at initialization and whenever a changein parameters such as, for example, number and/or position of zones, isrequired when the user selects a different scanning pattern. Thefront-end system controller 28 also provides buffering and flow controlfunctions, as data from the beamformer is sent to the host via abandwidth-constrained link, to prevent data loss.

The host computer 20 includes a keyboard/mouse controller 38, and adisplay controller 42 which interfaces with a display or recordingdevice 44. A graphical user interface described in co-pending U.S.patent application Ser. No. 09/822,764 entitled “Unitary OperatorControl for Ultrasonic Imaging Graphical User Interface,” by MichaelBrodsky, the entire contents of which are being incorporated herein byreference, may be used in a preferred embodiment of the presentinvention.

The host computer further includes a processing unit such asmicroprocessor 36. In a preferred embodiment of the ultrasound imagingsystem in accordance with the present invention the microprocessor 36includes on-chip parallel processing elements. In a preferredembodiment, the parallel processing elements may include a multiplierand an adder. In another preferred embodiment, the processing elementsmay include computing components, memories, logic and control circuits.Depending on the complexity of the design, the parallel processingelements can execute either SIMD or Multiple Instruction Multiple Data(MIMD) instructions.

Further, the host computer includes a memory unit 40 that is connectedto the microprocessor 36 and has a sequence of instructions storedtherein to cause the microprocessor 36 to provide the functions of downconversion, scan conversion, M-mode, and Doppler processing whichincludes color flow imaging, power Doppler and spectral Doppler, and anypost-signal processing. The down conversion or mixing of sampled analogdata may be accomplished by first multiplying the sampled data by acomplex value and then filtering the data to reject images that havebeen mixed to nearby frequencies. The outputs of this down-conversionprocessing are available for subsequent display or Doppler processing.

The scan conversion function converts the digitized signal data from thebeamforming circuitry 32 from polar coordinates (r,θ) to rectangularcoordinates (x,y). After the conversion, the rectangular coordinate datacan be forwarded for optional post signal processing where it isformatted for display on the display 44 or for compression in a videocompression circuit. Scan conversion and beamforming and associatedinterfaces are described in U.S. Pat. No. 6,248,073 to Jeffrey M.Gilbert et al., issued on Jun. 19, 2001, entitled “Ultrasound ScanConversion with Spatial Dithering,” the entire contents of which arebeing incorporated herein by reference.

The Doppler processing (CFI, PD, spectral Doppler) is used to imagetarget tissue 12 such as flowing blood. In a preferred embodiment, withpulsed Doppler processing, a color flow map is generated. In a preferredembodiment, the CPI, PD, Spectral Doppler computation can be carried outin software running on the host processor. Parallel computation unitssuch as those in the Intel® Pentium® and Pentium® III's MMX™coprocessors allow rapid computation of the required functions. Forparallel processing computation, a plurality of microprocessors arelinked together and are able to work on different parts of a computationsimultaneously. In another preferred embodiment, digital SignalProcessor (DSP) can also be used to perform the task. Such arrangementpermits flexibility in changing digital signal processing algorithms andtransmitting signals to achieve the best performance as region ofinterest is changed.

Single Instruction Multiple Data (SIMD) parallel processors allow onemicro-instruction to operate at the same time on multiple data items toaccelerate software processing and thus performance. One chip providescentral coordination in the SIMD parallel processing computer.Currently, SIMD allows the packing of four single precision 32-bitfloating point values into a 128-bit register. These new data registersenable the processing of data elements in parallel. Because eachregister can hold more than one data element, the processor can processmore than one data element simultaneously. In a preferred embodiment ofthe present invention, all the data is organized efficiently to use SIMDoperations. In a particular embodiment, Multiple Instruction MultipleData (MIMD) parallel processors may be used, which include a pluralityof processors. Each processor can run different parts of the sameexecutable instruction set and execute these instructions on differentdata. This particular embodiment employing MIMD may be more flexiblethan the embodiment utilizing SIMD, however may be more expensive. Allthe kernel functions such as demodulation, Gauss match filtering,Butterworth high pass filtering, auto-correlation calculation,phase-shift calculation, frame averaging, color-averaging, spatialdomain low-pass filtering, and scan conversion interpolation areimplemented with SIMD or SIMD. The Doppler processing results in theprocessed data being scan converted wherein the polar coordinates of thedata are translated to rectangular coordinates suitable for display orvideo compression.

The control circuit, preferably in the form of a microprocessor 36inside of a personal computer (e.g., desktop, laptop, palmtop), controlsthe high-level operation of the ultrasound imaging system 10. Themicroprocessor 36 or a DSP initializes delay and scan conversion memory.The control circuit 36 controls the differential delays introduced inthe beamforming circuitry 32 via the memory 26.

The microprocessor 36 also controls the memory 40 which stores data. Itis understood that the memory 40 can be a single memory or can bemultiple memory circuits. The microprocessor 36 also interfaces with thepost signal processing functional instructions and the displaycontroller 44 to control their individual functions. The displaycontroller 44 may compress data to permit transmission of the image datato remote stations for display and analysis via a transmission channel.The transmission channel can be a modem or wireless cellularcommunication channel or other known communication method.

The preferred embodiments of the ultrasonic imaging system address themain problem of performing Doppler processing by software whichtypically is the speed for processing in real time. FIG. 2 illustrates apreferred embodiment of a method 100 for Doppler processing inaccordance with the present invention.

The method 100 includes mapping or vectorizing of the serial input RFdata for parallel computation per step 102. Further details of the datamapping process are described in FIG. 3 which diagrammaticallyillustrates a preferred embodiment of a method 150 for data mapping forparallel processing. An input data steam 152 is mapped into an inputvector representation. The vectors 154, 156 are sequentially provided tothe microprocessor's 36 parallel processing elements in which aninstruction can be executed that allows all data within the vectors 154,156 to be operated on in parallel. The dimension of the vector P isdetermined by the hardware constraints of the microprocessor and mayequal the number of parallel processors. For example, current MMX™technology allows a vector with four dimensions, thus limiting theprocessing of four dimension vectors in parallel. The dimensions of thevector representation are not limited to the current availabletechnology and can accommodate increases in the dimension of the vectorrepresentation.

The method 100 includes the step 104 of smoothing the RF data to enhancethe signal to noise ratio (SNR), for example, by curve fittingtechniques. The RF data is then demodulated to generated IQ data, whereI represents in-phase and Q represents quadrature samples in step 106.Demodulation may be performed, but is not limited to, by a rectangularfilter.

In a preferred embodiment, Doppler processing is operated on eachsub-segment instead of each single sample. This method saves processingtime because the number of sub-segments is much less than number ofsamples. This is important for clinical use of software-based Dopplerproducts, where a real-time system is critical. Further, by processingsub-segments the sensitivity to flows increases. The data processed pera sequence of instructions in a preferred embodiment represents theaverage over a segment, which has higher signal to noise ratio.

The use of data segments may however cause lower spatial resolution. Toreduce such an effect, the signal sequence is not divided as individualsegments, instead each adjoining segment is overlapped with each other.In other words, the center distance or down-sample space between eachsegment is less than the length of segments. According to the Nyquistsampling theorem, if such a distance is less or equal to the halfsegment length, the spatial resolution can be recovered by properinterpolation methods. For example, if there exists a demodulated IQdata sequence y₀, y₁, y₂, y₃, y₄, y₅, y₆, y₇, the segment length as fourand down-sample space as two, a down-sampled data sequence z₀, z₁, z₂,is generated where z₀ is the average of y₀, y₁, y₂, y₃, z₁ is average ofy₂, y₃, y₄, y₅, z₂ is the average of y₄, y₅, y₆, y₇. Mathematically, theaverage over each segment can be expressed as

$\begin{matrix}{{z(t)} = {\frac{1}{L}{Re}\mspace{11mu} {{ct}\left( \frac{t}{L} \right)}*{y(t)}}} & (19)\end{matrix}$

where y(t) is demodulated complex IQ data,

$\begin{matrix}{{{Re}\mspace{11mu} {{ct}(t)}} = \left\{ \begin{matrix}1 & {{- 0.5} \leq t \leq 0.5} \\0 & {Otherwise}\end{matrix} \right.} & (20)\end{matrix}$

If the down-sample space is Δt≦L/2, then the down-sampled data can beexpressed as

z _(s)(k)=z(kΔt) where k=0,1,2, . . .  (21)

The method 100 further includes the step 108 of filtering thedown-samples using, for example, without limitation, the Gauss Matchfilter. A match filter is used to maximize the signal-to-noise ratio.The signal for Doppler processing is generated by performing quadraturedemodulation with the emitted frequency (ω₀) and then Gauss Matchfiltering the complex signal.

y(t)=Gauss(t)*[rf(t)exp(jω ₀ t)]  (22)

where rf(t) is the raw RF data. To save the calculation time, matchfiltering is only performed on down-sampled samples. From (19), itfollows

z(t)=Rect(t)*Gauss(t)*[rf(i)exp(jω ₀ t)]  (23)

Equation (23) can be also expressed as

z(t)=Gauss(t)*x(t)  (24)

x(t)=Rect(t)*[rf(t)exp(jω ₀ t)]  (25)

Thus, the down-sampled data can be expressed as

$\begin{matrix}{{z_{s}(k)} = {{z\left( {k\; \Delta \; t} \right)} = {\int_{- \infty}^{+ \infty}{{{Gauss}(\tau)} \times \left( {{k\; \Delta \; t} - \tau} \right)\ {\tau}}}}} & (26)\end{matrix}$

The method 100 further includes the step 110 of calculating the poweralong the sample line before a high pass filter or wall filter. Per step112, stationary signals are filtered out by a high pass filter, forexample, but not limited to, a Butterworth filter. The method 100 thenproceeds to step 114 wherein the auto-correlation function R(l) of thesignal is calculated.

Per step 116, the phase shift of the auto-correlation phase is thencalculated efficiently. As described hereinbefore, the mean velocity canbe determined by the mean angular frequency

$\begin{matrix}{\overset{\_}{\varpi} = \frac{\int_{- \infty}^{+ \infty}{\varpi \; P\; (\varpi)\ {\varpi}}}{\int_{- \infty}^{+ \infty}{P\; (\varpi)\ {\varpi}}}} & (27)\end{matrix}$

where P(ω) is the power density spectrum of Doppler signal. The meanblood flow velocity v can then be estimated by the following equation

$\overset{\_}{v} = {\frac{\varpi}{\omega_{0}}\frac{c}{2\mspace{11mu} \cos \mspace{11mu} \theta}}$

where c is the velocity of sound and θ the angle between the sound beamand the blood flow vector. It has been shown that:

$\begin{matrix}{\overset{\_}{\varpi} = {{{\overset{.}{\varphi}(0)} \approx \frac{{\varphi (1)} - {\varphi \left( {- 1} \right)}}{2}} = \varphi}} & (28)\end{matrix}$

Generally, φ(1) can be determined by either of the following methods

$\begin{matrix}{{\varphi (1)} = {\arctan \left( \frac{{Im}\left\{ {R(1)} \right\}}{{Re}\left\{ {R(1)} \right\}} \right)}} & (29) \\{{\varphi (1)} = {\arcsin \left( \frac{{Im}\left\{ {R(1)} \right\}}{{R(1)}} \right)}} & (30) \\{{\varphi (1)} = {\arccos \left( \frac{{Re}\left\{ {R(1)} \right\}}{{R(1)}} \right)}} & (31)\end{matrix}$

All these operations are too time consuming to be implemented bysoftware. Besides, these functions are not monotonic in the intervalbetween −pi and +pi (−π, +π). Thus, not all values of the phase shiftcorresponding to the range of Doppler velocities according to theNyquist criterion are accounted for. In a preferred embodiment of theDoppler processing system in accordance with the present invention

${{sign}\left( {\varphi (1)} \right)}{\sin^{2}\left( \frac{\varphi (1)}{2} \right)}$

represents die phase-shift, where

$\begin{matrix}{{{sign}(x)} = \left\{ {\begin{matrix}1 & {x > 0} \\0 & {x = 0} \\{- 1} & {x < 0}\end{matrix}{{Sign}\left( {\varphi (1)} \right)}{\sin^{2}\left( \frac{\varphi (1)}{2} \right)}} \right.} & (32)\end{matrix}$

is a monotonic function of φ(1) in the interval (−π, +π). In otherwords, every value of

${{sign}\left( {\varphi (1)} \right)}{\sin^{2}\left( \frac{\varphi (1)}{2} \right)}$

uniquely defines a φ(1) in the interval (−π, +π) and vice versa.Calculation of the sin² function avoids the use of a square rootoperation, which is computationally intensive.As we know

$\begin{matrix}{{{sign}\left( {\varphi (1)} \right)} = {{{sign}\left( {\sin \; \left( {\varphi (1)} \right)} \right)} = {{sign}\left( {{Im}\left\{ {R(1)} \right\}} \right)}}} & (33) \\{{{\sin^{2}\left( \frac{\varphi (1)}{2} \right)} = {\frac{1 - {\cos \left( {\varphi (1)} \right)}}{2} = \frac{1 - \frac{{Re}\left\{ {R(1)} \right\}}{{R(1)}}}{2}}}{So}} & (34) \\{{{{sign}\left( {\varphi (1)} \right)}{\sin^{2}\left( \frac{\varphi (1)}{2} \right)}} = {\frac{1}{2}{{{sign}\left( {{Im}\left\{ {R(1)} \right\}} \right)}\left\lbrack {1 - \frac{{Re}\left\{ {R(1)} \right\}}{{R(1)}}} \right\rbrack}}} & (35)\end{matrix}$

Similarly, the angle phi φ(1) may be also calculated by

$\begin{matrix}{{\tan \left( \frac{\varphi (1)}{2} \right)} = {\frac{\sin \mspace{11mu} {\varphi (1)}}{1 + {\cos \left( {\varphi (1)} \right)}} = \frac{\frac{{Im}\left\{ {R(1)} \right\}}{{R(1)}}}{1 + \frac{{Re}\left\{ {R(1)} \right\}}{{R(1)}}}}} & \left( {35a} \right)\end{matrix}$

The method 100 further includes the step 118 wherein the echo signalsare rejected by color priority. In a preferred embodiment, if the powerbefore the high pass filter is larger than a predetermined threshold,the phase shift is set to zero. Per step 120 the phase shift is smoothedby, but is not limited to, a spatial low-pass filter. The phase-shift isaveraged with previous frames such as, for example, the previous twoframes per step 122. The phase shift is converted to a color index perstep 124 to obtain a curve like the one represented by the phase shiftcalculation in equations 35 and 35a. In a preferred embodiment, a lookup table may be used, but is not limited to, in order to convert thephase shift to a color index. Color averaging is then performed per step126. Per step 128 the phase shift represented in scan coordinates istransformed or converted to raster coordinates using, for example, butis not limited to, interpolation methods. The resultant color images aresuperpositioned on the B-mode image in step 130.

FIG. 4A illustrates a graphical representation 170 of the Doppler phaseshift calculations, in particular for the following equations 10 and 15described hereinbefore. The graphical representation of equation 15 inaccordance with a preferred embodiment, is a monotonic function in theinterval between −pi and pi which spans the range of Doppler velocitiesaccording due to the Nyquist criterion. This can be expressed by simplemathematical operations, as shown in the right hand side of equation 15.

$\begin{matrix}{{\varphi (1)} = {\arcsin \left( \frac{{Im}\left\{ {R(1)} \right\}}{{R(1)}} \right)}} & (10) \\{{{{sign}\left( {\varphi (1)} \right)}{\sin^{2}\left( \frac{\varphi (1)}{2} \right)}} = {\frac{1}{2}{{{sign}\left( {{Im}\left\{ {R(1)} \right\}} \right)}\left\lbrack {1 - \frac{{Re}\left\{ {R(1)} \right\}}{{R(1)}}} \right\rbrack}}} & (15)\end{matrix}$

FIG. 4B illustrates a graphical representation 190 of equations 9 and15a.

$\begin{matrix}{{\varphi (1)} = {\arctan \left( \frac{{Im}\left\{ {R(1)} \right\}}{{Re}\left\{ {R(1)} \right\}} \right)}} & (9) \\{{\tan \left( \frac{\varphi (1)}{2} \right)} = {\frac{\sin \mspace{11mu} {\varphi (1)}}{1 + {\cos \left( {\varphi (1)} \right)}} = \frac{\frac{{Im}\left\{ {R(1)} \right\}}{{R(1)}}}{1 + \frac{{Re}\left\{ {R(1)} \right\}}{{R(1)}}}}} & \left( {15a} \right)\end{matrix}$

Equation 15a, is the graphical representation of a preferred embodiment,and is a monotonic function in the interval between −pi and pi whichspans the range of Doppler velocities according to the Nyquistcriterion, and can be expressed by simple mathematical operations, asshown in the right hand side of equation 15a.

It should be noted that an operating environment for the system 10includes a processing system with at least one high speed processingunit and a memory system. In accordance with the practices of personsskilled in the art of computer programming, the present invention isdescribed with reference to acts and symbolic representations ofoperations or instructions that are performed by the processing system,unless indicated otherwise. Such acts and operations or instructions aresometimes referred to as being “computer-executed”, or “processing unitexecuted.”

It will be appreciated that the acts and symbolically representedoperations or instructions include the manipulation of electricalsignals by the processing unit. An electrical system with data bitscauses a resulting transformation or reduction of the electrical signalrepresentation, and the maintenance of data bits at memory locations inthe memory system to thereby reconfigure or otherwise alter theprocessing unit's operation, as well as other processing of signals. Thememory locations where data bits are maintained are physical locationsthat have particular electrical, magnetic, optical, or organicproperties corresponding to the data bits.

The data bits may also be maintained on a computer readable mediumincluding magnetic disks, optical disks, organic disks, and any othervolatile or non-volatile mass storage system readable by the processingunit. The computer readable medium includes cooperating orinterconnected computer readable media, which exist exclusively on theprocessing system or is distributed among multiple interconnectedprocessing systems that may be local or remote to the processing system.

In preferred embodiments, the ultrasound imaging system includes colorand power Doppler modes to detect the presence, direction and relativevelocity of blood flow by assigning color-coded information to theseparameters. The color is depicted in a region of interest that isoverlaid on to the B-mode image. All forms of ultrasound-based imagingof red blood cells are derived from the echo signal that is received inresponse to the transmitted signal. The primary characteristics of thissignal are its frequency and its amplitude (or power). The frequencyshift is determined by the movement of the red blood cells. Theamplitude is dependent on the amount of blood in motion present withinthe volume that is sampled by the ultrasound beam.

In a preferred embodiment, as illustrated in FIG. 5, color Dopplerallows flow velocity information to be detected over a portion of theB-Mode image. The mean Doppler shift is then displayed against a grayscale representation of the structures. Color Doppler images displayblood flow by mapping the frequency shift (velocity) of blood cells.Flow towards the transducer is assigned shades of red, and flow awayfrom the transducer is displayed in shades of blue. Alternateembodiments can use different color contrasts. Higher frequencies aredisplayed in lighter colors and lower frequencies in darker colors. Forexample, the proximal carotid artery is normally displayed in hues ofbright red and orange because the flow is toward the transducer and thefrequency (velocity) of flow in this artery is relatively high. Bycomparison, the flow in the jugular vein is displayed as blue because itflows away from the transducer.

The color Doppler mode provides, in preferred embodiments, fourteenexamination types to provide presets depending upon the probes used. Therange of WF includes 1% to 25% of pulse repetition frequency (PRF), andsteering angles include, for example, +/−15°, +/−20°, and 0 for flatlinear probes. Further, maps in red/blue based on velocity of blood flowand high spatial resolution versus high frame rate for certain probes isalso provided. In a preferred embodiment, as illustrated in FIGS. 6A and6B the color Doppler (CD)-mode for image display is selected by choosingthe CD-mode button from a image mode bar, or from a Modes menu. Toselect the CD image control setting, the user selects the CD tab 242 atthe bottom of the image control bar. The scan area controls the size ofthe region of interest. Further, in preferred embodiments the Dopplersystems use a wall filter to reduce and preferably eliminate unwantedlow frequency signals from the display. The color controls 244 a-d canbe adjusted to increase the quality of an image.

FIGS. 7A-7D illustrate the controls and the image sizes available in acolor Doppler mode in accordance with a preferred embodiment. The sizeof the scan area is one of the controls the user manipulates to affectthe frame rate. The smaller the scan area, the faster the frame rate.Alternatively, the larger the scan area, the slower the frame rate. Forexample, for cardiac or arterial applications, a small scan area can beused to accurately visualize the flow dynamics. A medium or large scancan also be used for applications where the blood flow dynamics do notchange rapidly over time, or if the user wants to get a larger overallview of the blood flow.

In a preferred embodiment, as shown in FIG. 8A, the ultrasound systemprovides for adjustment of the pulse repetition frequency (PRF). The PRFadjusts the range of the velocity display shown on a color bar.Decreasing the PRF values improves the display of slow blood flow. Themaximum value for the PRF is dependent on the specific probe that isbeing used and how deep the region of interest is. The PRF can be sethigh enough to prevent aliasing, and low enough to provide adequatedetection of low flow. It may be necessary to vary the PRF during anexamination, depending on the speed of the blood flow, and/or if apathology is present. If the PRF is set too high, low frequency shiftscaused by low velocity flow may not be shown. In preferred embodiments,the PRF is set higher for cardiac and arterial applications than it isfor venous or small parts application.

In a preferred embodiment, Doppler systems use a wall filter toeliminate unwanted low frequency signals from the display. Raising thewall filter reduces the display of low velocity tissue motion. Asillustrated in FIG. 8B, the ultrasound system provides for an adjustmentof a wall filter. As the wall filter is lowered, more information isdisplayed, however, more wall tissue motion can also be displayed. Thewall filter can range from 1% to 25% of the PRE (in increments of 2, 10,25, 50 or 100 depending on the wall filter value). The wall filter isset high enough so that color Doppler flash artifacts from tissue orwall motion are not displayed, but low enough to display slow flow. Ifthe wall filter is set too high, slower flow may not be seen. The wallfilter is set higher for applications where there is significant tissuemotion (such as a cardiac application), or in instances where the probeis moved rapidly while color Doppler is on. It may be set lower forsmall parts or instances where flow is slow but there is not much tissuemotion.

A preferred embodiment also provides for the adjustment of a steeringangle as shown in FIG. 8C. The steering angle is used only on flatlinear arrays, and allows adjustments for optimal angle between beam andflow. The angle range is, but not limited to, for example, 0, +/−15°,+/−20°. Doppler angle adjustment is important in obtaining adequatesignals and in interpreting the color Doppler examination. Changes inthe angle of the Doppler beam to the path of flow causes changes in thecolor Doppler display. When using color Doppler, the user has to beaware of the Doppler angle to flow. At a 90 degree angle to flow, anabsent or confusing color pattern is displayed (even when the flow isnormal). An adequate Doppler angle to flow is required in order toobtain useful color Doppler information. In most instances, the lowerthe angle of the color Doppler beam to flow, the better the receivedsignal. Electronic steering is useful in the embodiments where the flowis at a poor angle to the color Doppler beam. However, in manyembodiments it is also necessary to use the “heel to toe” technique,which involves using pressure on one end of the probe or the other toimprove the Doppler angle to flow. It should be noted that curved linearand phased array probes do not have the capability of electronicsteering and depending on the clinical situation, may require the use ofthe “heel and toe” technique to obtain an adequate angle to flow. Ascolor Doppler is angle dependent, a weak flow signal results if thevessel is approximately at 90 degrees to the beam.

Another preferred embodiment includes the provision of color inversion.When the user chooses the Color Invert box illustrated in FIG. 8D, thecolor Doppler reference bar is inverted as well as the color within theregion of interest. The color reference bar is divided by a zerobaseline. To invert the color Doppler reference bar, the user can clickon the Color Invert box. The Play button is used to see the Color Inverteffect on a frozen image. Conventionally, the color red is assigned topositive frequency shifts (flow toward the probe), and blue is assignedto negative frequency shifts (or flow away from the probe). However,this color assignment can be reversed at the user's discretion byselecting Invert. Whether or not the user has inverted the display, flowtoward the probe is assigned the colors of the top half of the colorbar, and flow away from the probe is assigned the colors of the bottomhalf of the color bar. FIGS. 9A and 9B illustrate the inverted andnon-inverted reference bar in the color Doppler mode. Invert may be usedwhen scanning the internal carotid artery (ICA), for example. Ingeneral, flow in this vessel goes away from the probe. If Invert isenabled, the ICA flow is displayed in shades of red. The color bar thendisplays shades of blue on the top half, and a shade of red on thebottom.

Preferred embodiments also allow for the adjustment of the color gainwhich increases or decreases the amplification of the returning signaldisplayed or being played. The color gain adjustment 352 is illustratedin FIG. 10. The color gain can be increased whenever the fill of colorwithin a vessel is inadequate, and can be decreased whenever anunacceptable amount of color is seen outside of a vessel.

Preferred embodiments also accommodate the adjusting of the colorpriority of the image which defines the amount of color displayed overbright echoes and helps confine color within the vessel walls. Thisadjustment in color priority 354 is illustrated in FIG. 10. This controlaffects the level at which color information overwrites the B-modeinformation. If the user needs to see more flow in an area of somesignificant B-mode brightness, they increase the color priority. If auser wants to better contain the display of flow within the vessel(s),they decrease the color priority. If, however, the color priority is setall the way to the left, no color is displayed.

A preferred embodiment also accommodates for adjusting the colorpersistence as illustrated in FIG. 10. The color persistence adjustment356 averages frames of color. Increasing the persistence causes thedisplay of flow to persist on the 2D image. Decreasing the persistenceallows better detection of short duration jets, and provides a basis forbetter flow/no flow decisions. Adjusting color persistence also producesbetter vessel contour depiction. To adjust the color persistence for ascan, in a preferred embodiment, the slider is moved to the right orleft to achieve the desired image. When color persistence is set high,the saved image (single frame) may not look exactly the same as when theimage is saved. The user receives a warning when this occurs; however,this does not occur when exporting images.

In a preferred embodiment, as illustrated in FIG. 10, the color baseline358 can also be adjusted. The baseline refers to the zero baselinewithin the color Doppler window. Adjusting this control moves the zerobaseline up or down. To adjust the baseline, the slider is moved to theright or left to achieve the desired range of degree. In general, it isnot necessary to adjust the color baseline. If the user does adjust it,moving the baseline down displays more positive flow and moving thebaseline up displays more negative flow in an embodiment. When the useradjusts the color baseline, they are able to display more forward orreversed flow. This adjustment can be used to prevent aliasing in eitherdirection.

Further embodiments include an adjustment for high frame rate 360 shownin FIG. 10. This control adjusts the line density within the Dopplerregion of interest. The user can choose between an image with higherline density resulting in better spatial resolution or lower linedensity resulting in a better frame rate. The high frame rate is usedwhen the flow rate is high such as in cardiac or certain arterialapplications. This control is available with certain embodiments of theprobes. When the user selects high spatial resolution, line density isincreased. However, when the user selects high frame rate, line densityis decreased.

Another preferred embodiment of the present invention includes adirectional power Doppler mode which can be viewed as a combination ofboth conventional power Doppler and color Doppler modes. It provides thesame increased sensitivity as conventional power Doppler as well asdirectional information derived from color Doppler. Directional powerDoppler does not provide an estimate of the frequency (velocity) ofblood flow. The color palette is proportional to the strength of theDoppler signal. The mode allows the user to achieve good image qualityof deep arteries and other tissue. One also has the option to apply ahigh frame rate or high resolution to control the quality of the scan.The benefits of directional power Doppler include examination types toprovide presets (dependent upon certain probes), Doppler steering anglesof (+/−15°, =/−20°, 0°, for linear probes), without limitation, range ofWF and increments (2 to PRF/4 Hz) and increased Doppler sensitivityperformance.

As illustrated in FIGS. 11A and 11B, the DirPwr-node 382 is selected andfor image display, the DirPwr-mode button from the image mode bar can bechosen, or from the Modes menu. To select the DirPwr image controlsetting, the user selects the DirPwr tab at the bottom of the imagecontrol bar. Similar to the color Doppler mode, the scan area controlsthe size of the region of interest. The Doppler systems use the wallfilter to eliminate unwanted low frequency and provides color controladjustments.

As illustrated hereinbefore, the scan can be adjusted even in thedirectional power Doppler mode. The size of the scan area is one of thecontrols that is used to affect the frame rate. The smaller the scanarea, the faster the frame rate. Alternatively, the larger the scanareas, the slower the frame rate. For cardiac or arterial applications,a small scan area is used to accurately visualize the flow dynamics. Amedium or large scan can also be used for applications where the bloodflow dynamics do not change rapidly over time, or if the user wishes toget a larger overall view of the blood flow. The region of interest canbe defined by adjusting the scan area. The scan area options range fromsmall, medium to large.

In the directional power Doppler mode, the PRF can be set high enough toprevent aliasing, and low enough to provide adequate detection of lowflow. It may be necessary to vary the PRF during one examinationdepending on the speed of the blood flow, and/or if the pathology ispresent. If the PRF is set too high, signals caused by slow flow stateswith few red blood cells may not be shown. In an embodiment, the PRF isset higher for high flow states than it is for low flow states. The PRFcan be set high enough to prevent aliasing, and low enough to provideadequate detection of low flow. It may be necessary to vary the PRFduring one examination, depending on the speed of the blood flow, and/orif the pathology is present. If flow is weak or slow it is displayed indarker shades. If flow is strong or fast it is shown in brighter shades.Directional power Doppler is somewhat angle dependent. Good Dopplerangles can be maintained.

Further, Doppler systems use a wall filter to eliminate unwanted lowfrequency signals from the display. Raising the wall filter reduces thedisplay of low velocity tissue motion. As the wall filter is lowered,more information is displayed, however, more wall tissue motion is alsodisplayed. The wall filter can be set high enough so that Doppler flashartifacts from tissue or wall motion are not displayed, but can be setlow enough to display slow flow. If the wall filter is set too high,slower flow may not be seen. The wall filter is set higher forapplications where there is significant tissue motion, or in instanceswhere the probe is moved rapidly while directional power Doppler isenabled. It may be set lower for small parts or instances where flow isweak but there is not much tissue motion. The wall filter range isdependent on which probes the user is using as well as the PRF setting.Wall filter range is from 1% to 25% of the PRF (in increments of 2, 10,25, 50 or 100 depending on the wall filter value).

In addition, the steering angle adjusts the optimal angle to flow in thedirection power Doppler mode. The steering angle is used only on lineararrays, and allows adjustments for optimal angle between beam and flow.The angle range is without limitation, 0, +/−15°, +/−20°. The steeringangle adjusts the optimal angle to flow. Doppler angle adjustment isimportant in obtaining adequate signals and in interpreting the colorDoppler examination. Changes in the angle of the Doppler beam to thepath of flow causes changes in the color Doppler display. When usingcolor Doppler the user has to be aware of the Doppler angle to flow. Ata 90 degree angle to flow, an absent or confusing color pattern isdisplayed (even when the flow is normal). An adequate Doppler angle toflow is required in order to obtain useful color Doppler information. Inmost instances, the lower the angle of the color Doppler beam to flow,the better the received signal. Electronic steering is available forflat linear array probes. Electronic steering is useful in preferredembodiments in those instances where the flow is at a poor angle to thecolor Doppler beam. However, in many instances it is also necessary touse the “heel and tow” technique, which involves using pressure on oneend of the probe or the other to improve the Doppler angle to flow.Curved linear and phased array probes do not have the capability ofelectronic steering and depending on the clinical situation, may requireyou to use the “heel and toe” technique to obtain an adequate angle toflow. To change the steering angle, the down arrow is clicked and thedesired setting is selected.

In an embodiment including directional power Doppler, the user canadjust the color gain which increases or decreases the amplification ofthe returning signal displayed or being played. Color gain can beincreased whenever the fill or color within a vessel is inadequate, andshould be decreased whenever an unacceptable amount of color is seenoutside of a vessel.

Further, in an embodiment including directional power Doppler, colorpriority can be adjusted. This control affects the level at which colorinformation overwrites the B-mode information. If the user needs to seemore flow in an area of some significant B-mode brightness, the colorpriority is increased. If the user wants to better contain the displayof flow Within the vessel(s), they decrease the color priority. Toadjust the color priority the slider is moved to the right or left toachieve the desired range. Raising the priority displays color onbrighter structures. Lowering priority increases containment of thedisplay of flow to within the vessels. If, however, color priority isset all the way to the left, no color will be displayed.

The color persistence can be adjusted which averages frames of color.Increasing the persistence causes the display of flow to persist on the2D image. Decreasing the persistence allows better detection ofshort-duration jets, and provides a basis for better flow/no flowdecisions. Adjusting color persistence also produces better vesselcontour depiction.

Further, in the preferred embodiment it is not necessary to adjust thecolor baseline. If the user does not wish to adjust it, moving thebaseline down displays more positive flow and moving the baseline updisplays more negative flow. When you adjust the color baseline, you areable to display more forward or reversed flow. This adjustment can beused to prevent aliasing in either one direction.

The preferred embodiment includes a high frame rate adjustment. Thiscontrol adjusts the line density within the Doppler region of interest.The user can choose between one image with higher line density resultingin better spatial resolution or lower line density resulting in betterframe rate. High frame rate is used when the flow rate is high such asin cardiac or certain arterial applications. This control is onlyavailable on certain probes.

Conventional power Doppler in accordance with a preferred embodimentimages blood flow by displaying the density of red blood cells, asopposed to their velocity. Large amplitude signals are assigned a brighthue and weak signals are assigned a dim hue. All flow is displayed inshades of the same color. No directional information is provided in thisembodiment.

In accordance with another embodiment, the ultrasound imaging systemincludes a power Doppler mode. The sensitivity of power Doppler isgreater than color Doppler. Amplitude estimation in this mode is lessnoisy than a mean frequency estimate, therefore more real signal isdetected and displayed with the power Doppler mode. Power Doppler isalso less angle dependent than Color Doppler because of this increase insensitivity, and does not alias. All flow is displayed in shades of thesame color, however, no directional information is provided. Thebenefits of power Doppler include it being more sensitive to low flowthan color or direction power Doppler, it is the preferred mode to showperfusion and contour of vessel lumen and large amplitude signals areassigned a bright hue and weak signals are assigned a dim hue. Forexample, the jugular vein is shown in brighter colors than the carotidartery because the vein contains more red blood cells at any given timethan does the artery.

To select the Pwr-mode for image display, the user select a Pwr-modebutton from the image mode bar, or from the Modes menu as illustrated inFIGS. 12A and 12B. To select the Pwr image control setting, the userselects the Pwr tab 412 at the bottom of the image control bar.Increasing the PRF increases the range of the display. Lower PRF valuesimprove the display of slow flow. PRF is dependent on the specific probethat is being used.

In the embodiment including power Doppler mode, the size of the scanarea is one of the controls the user has access to that most affects theframe. The smaller the scan area, the faster the frame rate.Alternatively, the larger the scan area, the slower the frame rate. Forcardiac or arterial applications, a small scan area is used toaccurately visualize the flow dynamics. A medium or large scan can alsobe used for applications where the blood flow dynamics do not changerapidly over time, or if the user wishes to get a larger overall view ofthe blood flow. The region of interest can be defined by adjusting thescan area. The scan area options range from small, medium to large aspreviously described.

In the power Doppler mode, the PRF (Pulse Repetition Frequency) can below enough to provide adequate detection of low amplitude flow. It maybe necessary to vary the PRF during one examination depending on thespeed of the blood flow, and/or if the pathology is present. If the PRFis set too high, signals caused by slow flow states with few red bloodcells may not be shown. The PRF is set higher for high flow states thanit is for flow states in most preferred embodiments. To increase thePRF, the right arrow is selected. To decrease the PRF, the left arrow ina control is selected. The PRF values range from 100 Hz to 15 kHz anddepend upon the probe being used. If there are few red blood cells beingsampled, flow is shown in the darker shades of gold. If there are manyred blood cells being sampled, flow is displayed in a brighter shade ofgold. Power Doppler is the least angle dependent Doppler mode. However,it is good practice to maintain reasonable Doppler angles while usingthis mode.

In an embodiment including power Doppler mode, a wall filter can be sethigh enough so that power Doppler flash artifacts from tissue or wallmotion are not displayed. However, the wall filter can be set low enoughto display low amplitude signals. If the wall filter is set too high,low amplitude signals may not be seen. The wall filter is set higher forapplications where there is significant tissue motion, or in instanceswhere the probe is moved rapidly while power Doppler is on. It may beset lower for small parts or instances where the amount of blood flow issmall and there is not much tissue motion.

With respect to adjusting the steering angle, power Doppler is lessangle dependent than Color Doppler, but many of the same rules stillapply. Changes in the angle of the Doppler beam to the path of flowcauses changes in the power Doppler display. At 90 degrees to flow, aweak signal may be displayed even when the flow is normal. The steeringangle adjusts the optimal angle to flow. To increase the steering angle,the down arrow to the desired setting is selected. The angle range is,for example, −0, +/−15, +/−20 without limitation. If an adequate angleis used, the Doppler angle and the quality of the received signal isenhanced. The more red cells being sampled the stronger the receivedsignal. Further, attenuation plays a role in the strength of thedisplayed signal. The closer the area of interest is to the probe, theless the attenuation and the stronger the received signal. In apreferred embodiment, electronic steering is available for flat lineararray probes. Electronic steering is useful in those instances where theflow is at a poor angle to the power Doppler beam. In many instances, itis also necessary to use the “heel and toe” technique, which involvesusing pressure on one end of the probe or the other to improve theDoppler angle to flow. Curved linear and phased array probes do not havethe capability of electronic steering and depending upon the clinicalsituation, may require you to use the “heel and toe” technique to obtainan adequate angle to flow.

Embodiments including power Doppler also provide for adjusting colorgain. Color gain can be increased whenever the fill of color within avessel is inadequate, and should be decreased whenever an unacceptableamount of color is seen outside of a vessel. The user can adjust thecolor gain which modifies the degree of sensitivity to receive signalsand increases or decreases the amplification of the returning signaldisplayed/played. To adjust the color gain, the slider is moved to theright or left to achieve the desired gain. The color gain maximumamplitude is, for example, 1 to 30 units.

In a preferred embodiment including a power Doppler mode, the colorpriority control affects the level at which color information overwritesthe B-mode information. If the user needs to see more flow in an area ofsome significant B-mode brightness, the color priority is increased. Ifthe user wants to better contain the display of flow within thevessel(s), they decrease the color priority. To adjust the colorpriority the slider is moved to the right or left to achieve the desiredrange. Raising the priority displays color on brighter structures.Lowering priority increases containment of the display of flow to withinthe vessels. The color priority maximum amplitude is, for example, inthe range of 1 to 30 units.

Preferred embodiments of the power Doppler mode include adjusting thecolor persistence which averages frames of color. Increasing thepersistence causes the display of flow to persist on the two-dimensionalimage. Decreasing the persistence allows better detection of shortduration jets, and provides a basis for better flow/no flow decisions.Adjusting color persistence also produces better vessel contourdepiction.

In an embodiment, the user has the option of selecting a high frame rateversus a high quality image. If the user wants a high quality image, thebutton for high spatial resolution is selected. If the user wants a highframe rate, the button for high frame rate is selected. When the userselects high spatial resolution, line density is increased. When theuser selects high frame rate, line density is decreased.

A preferred embodiment includes a pulsed wave Doppler mode, which is amode of scanning in which a series of pulses are used to study themotion of blood flow at a small region along a desired scanline, calledthe sample volume or sample gate. As illustrated in FIGS. 13A and 13Bthe x axis of the graph represents time while the y axis representsDoppler frequency shift. The real-time mixed mode is provided that showsthe B-mode scan area (FIG. 13A) along with a bottom window (FIG. 13B)that displays the pulsed Doppler waveform. The shift in frequencybetween successive ultrasound pulses caused mainly by moving red bloodcells, can be converted into velocity if an appropriate angle betweenthe insonating beam and blood flow is known. The strength of the signalappears as shades of gray, for example, in this spectral display. Thethickness of the spectral signal is indicative of laminar or turbulentflow (laminar flow typically shows a narrow band of blood flowinformation). In a preferred embodiment, pulsed wave Doppler and B-modeare shown together in a mixed mode display. This provides the user withthe capability of monitoring the exact location of the sample volume onthe B-mode image while doing pulsed wave Doppler. The benefits of pulsedwave Doppler include a plurality of examination types, for example,fourteen examination types to provide presets, the range of wall filteris between 1% to 25% of the Pulsed Repetition Frequency (PRF),measurements are available for all examination types including peaksystole (PS), end diastole (ED), Resistive Index (RI), andsystole-diastole ratio (PS/ED) and duplex imaging is available forreal-time display using B-mode and pulsed wave Doppler. The pulse waveDoppler (PWD) mode is selected for image display by selecting a PWD modebutton from the image mode bar, or from the modes menu. The PWD imagecontrol setting is selected using the PWD tab 452 at the bottom of theimage control bar.

In a preferred embodiment of a pulsed wave Doppler mode, there are threesweep speeds available. Further, adjustments to the velocity display areavailable which changes the unit of the vertical scale within the pulsedwave Doppler window between cm/s and KHz. The centimeter unit isavailable when the correction angle is situated between 0 and +/−70° ina preferred embodiment.

Preferred embodiments of the pulsed wave Doppler window includeadjustments to the PRF (Pulse Repetition Frequency) which adjusts thevelocity range of the display. The maximum value for PRF is dependent onthe specific probe that is being used and the location of the samplevolume. As PRF increases, the maximum Doppler shift that can bedisplayed without aliasing also increases. Lower PRF values improves thedisplay of slow blood flow. Increasing the PRF also increases thethermal index value.

In a preferred embodiment including the pulsed wave Doppler mode, a wallfilter (high pass frequency filter) is used to reduce and preferablyeliminate unwanted low frequency high-intensity signals (also known asclutter) from the display. Clutter can be caused by tissue motion or byrapid movement of the probe. A wall filter that is high enough to removeclutter is used but that is low enough to display spectral informationnear the baseline. The range of wall filter values is, for example,between 1% and 25% of the PRF value.

Preferred embodiment of the pulsed wave Doppler mode include adjustingthe steering angle. The steering angle is used only on flat lineararrays, and allows adjustment for optimal angle between beam and flow.The Doppler signal is weak when the angle between the insonating beamand blood flow approximates 90 degrees. Angles of 60° or less arerecommended for steering angle. The steering angle does not directlyaffect the calibration of the velocity scale.

A preferred embodiment also provides for adjusting an invert option. Theuser can invert the pulsed Doppler waveform. The Doppler scale isseparated by a zero baseline that extends across the width of thespectral display. The data above the baseline is classified as forwardflow. The data below the baseline is classified as reverse flow. Wheninverted, the reverse flow displays above the baseline and the forwardflow appears below the baseline. To invert the flow, an Invert box isselected.

A preferred embodiment also includes the provision of adjusting acorrecting angle in the pulsed wave Doppler mode. To obtain accuratevelocities, the user maintains Doppler angles of 60° or less. However,the user can employ higher values of correction angle, especially inperipheral vascular applications where the blood vessels are moreparallel to the face of the probe. In a preferred embodiment the maximumvalue for the correction angle is +/−70°. Velocity display in cm/s isshown only in the range between 0 and +/−70°. Above 70°, the error inthe velocity calculation is too large and the velocity scale isconverted automatically to frequency, independent of the correctionangle. However, the flow direction cursor is shown on the screen forreference. The correction angle control is active also on frozen images.To adjust the correction angle, the user selects the left and rightarrows in the Correction Angle field and clicks to the desired valuesetting. Or, in the alternative, the user can employ the quickadjustment key to easily set the angle correction between +/−60° degreesand 0 degrees.

Further, a preferred embodiment provides for adjusting sample volumesize and position in the pulsed wave Doppler mode. The user can adjustthe size of the pulsed wave Doppler region being examined by setting thesample volume size control. The lower the value, the narrower the samplesize used in the calculation of flow velocity contents. To adjust thesample volume size, the user selects the left and right arrows next tothe SV Size field to set to the desired value. The value range forsample volume size is, for example, 0.5 to 20 mm (in 0.5 mm increment).The adjustment is visible on the SV gate and Probe Info interface whenenabled. The position of the sample volume can be adjusted by using thetouch-pad control. Left-click on the sample volume (the line becomesgreen,) is selected to move it to the desired location, and left-clickis selected again to anchor it to make active. Alternatively, thismovement can be accomplished by using the keyboard arrow keys. Modifyingthe depth location of the sample volume affects the thermal index value.

Preferred embodiments of the pulsed wave Doppler mode include theadjustments for gain baseline and sound volume. The user can adjust thegain which increases or decreases the amplification of the returningsignal displayed/played. The gain can be adjusted so that the spectralwaveform is bright but not so high that the systolic window fills in, orother artifacts are created. The baseline refers to the zero baselinewithin the Pulsed Wave Doppler window. Adjusting this control moves thezero baseline up or down. When the user adjusts the baseline, they areable to display more forward or reversed flow, taking advantage of thefull scale available at that particular PRF value. This adjustment isvisible on the reference bar. Adjusting the sound volume control letsthe user define the volume of the pulsed wave Doppler. The sound volumeof the spectral signal can be adjusted to a comfortable level. If it istoo high, system noise may interfere with the sound produced by theblood flow.

It should be understood that the programs, processes, methods andsystems described herein are not related or limited to any particulartype of computer or network system (hardware or software), unlessindicated otherwise Various types of general purpose or specializedcomputer systems may be used with or perform operations in accordancewith the teachings described herein.

In view of the wide variety of embodiments to which the principles ofthe present invention can be applied, it should be understood that theillustrated embodiments are exemplary only, and should not be taken aslimiting the scope of the present invention. For example, the steps ofthe flow diagrams may be taken in sequences other than those described,and more or fewer elements may be used in the block diagrams. Whilevarious elements of the preferred embodiments have been described asbeing implemented in software, in other embodiments hardware or firmwareimplementations may alternatively be used, and vice-versa.

It will be apparent to those of ordinary skill in the art that methodsinvolved in the system and method for ultrasound imaging may be embodiedin a computer program product that includes a computer usable medium.For example, such a computer usable medium can include a readable memorydevice, such as, a hard drive device, a CD-ROM, a DVD-ROM, or a computerdiskette, having computer readable program code segments stored thereon.The computer readable medium can also include a communications ortransmission medium, such as, a bus or a communications link, eitheroptical, wired, or wireless having program code segments carried thereonas digital or analog data signals.

The claims should not be read as limited to the described order orelements unless stated to that effect. Therefore, all embodiments thatcome within the scope and spirit of the following claims and equivalentsthereto are claimed as the invention.

The invention claimed is:
 1. A portable ultrasound system for scanning aregion of interest with ultrasound energy to image the region ofinterest comprising: a battery powered, handheld processor housinghaving a display, a beamformer device and a system controller connectedto the beamformer device; a data processing system within the handheldprocessor housing, the data processing system having a processor thatreceives mapped data, the data processing system being programmed toperform a down conversion operation, a scan conversion operation and aDoppler processing operation that applies a programmable filter, thehandheld processor housing further including a memory in the processorhousing that is electrically connected to the data processing system,the memory having stored therein a sequence of executable instructionsthat are used to select a pulse repetition frequency from a plurality ofselectable values and to apply the programmable filter to ultrasoundimage data, the programmable filter including a wall filter having aplurality of selectable values, the sequence of instructions configuredto execute a calculation of an autocorrelation function and anautocorrelation phase such that the plurality of parallel processingunits generate Doppler data; and a graphical user interface configuredto operate on the display such that a user can adjust the programmablefilter with a first control window and adjust color priority with asecond control window to change a displayed color gain of an ultrasoundimage.
 2. The ultrasound system of claim 1 wherein the programmablefilter comprises a high pass filter.
 3. The ultrasound system of claim 2wherein the high pass filter comprises a Butterworth filter.
 4. Theultrasound system of claim 1 wherein the data processing system includesa parallel computing element having a multiplier and an adder.
 5. Theultrasound system of claim 1 wherein the data processing system executesat least one of Single Instruction Multiple Data instructions andMultiple Instruction Multiple Data instructions.
 6. The ultrasoundsystem of claim 1 wherein Doppler processing is executed with the dataprocessing system having a computation device including an MMX orfloating point processor.
 7. The ultrasound system of claim 1 whereinthe data processing system further comprises a directional power Dopplermode, a power Doppler mode and a pulsed wave Doppler mode.
 8. Theultrasound system of claim 1 further comprising an adjustable controlfor one of at least scan area selection, velocity display, steeringangles, color inversion, color gain, color priority, color persistence,color baseline and frame rate.
 9. The ultrasound system of claim 1wherein the wall filter has a selected value that is a function of thepulse repetition frequency.
 10. The ultrasound system of claim 1 whereinthe data processing system determines a tissue motion velocity.
 11. Theultrasound system of claim 1 wherein the data processing system comparesa Doppler signal to a threshold value such that if the Doppler signal islarger than the threshold value, a phase shift of the Doppler signal isset to zero.
 12. The ultrasound system of claim 11 wherein the thresholdvalue is selected from at least one of a B-mode value, a tissue motionvelocity and a high pass filtered Doppler velocity.
 13. The ultrasoundsystem of claim 1 wherein the graphical user interface further includesdisplay of a filter adjustment setting.
 14. A method for Dopplerprocessing of an ultrasound image on a portable ultrasound systemcomprising: processing ultrasound image data with a battery powered,handheld display device that includes a beamformer device, a systemcontroller to control beamformer device operation, a flat panel display,and a programmable data processing system having a processor configuredto perform a down conversion operation, a scan conversion operation anda Doppler processing operation, the portable ultrasound system includinga memory that is electrically connected to the programmable dataprocessing system, the memory having stored therein a sequence ofinstructions executed with the processor to use a pulse repetitionfrequency, to apply a programmable wall filter to image data and toexecute further instructions for a calculation of an autocorrelationfunction and calculation of an autocorrelation phase to perform Dopplerprocessing of the image data; using a graphical user interface operatingto select a pulse repetition frequency with a first control window shownon the display and select a wall filter value with a second controlwindow shown on the display; and performing a Doppler processingoperation using the processor of the programmable data processing systemto form an image of the region of interest.
 15. The method of claim 14further comprising performing Doppler processing including one of colorDoppler, directional power Doppler, power Doppler and pulsed waveDoppler.
 16. The method of claim 14 further comprising removing lowfrequency signals from the image data with the filter.
 17. The method ofclaim 14 wherein a color Doppler image is overlaid with a B-mode image.18. The method of claim 16 further comprising adjusting a wall filtervalue based on the pulse repetition frequency to display low velocitytissue motion, the wall value being between 1 percent and 25 percent ofthe pulse repetition frequency.
 19. The method of claim 14 furthercomprising filtering the image data with a Butterworth filter.
 20. Themethod of claim 14 further comprising adjusting the filter to imagecardiac motion.
 21. The method of claim 14 further comprisingdetermining a phase shift of the auto-correlation function to generate amonotonic function for all values of the phase shift corresponding to arange of Doppler velocities.
 22. The method of claim 14 wherein thesequence of instructions further comprises: filtering the data to removelow frequency signals; averaging a plurality of frames of data;converting phase shift data to an index; converting the phase shift datafrom scan to raster coordinates; and displaying a plurality of images.23. The method of claim 14 wherein the programmable data processingsystem includes at least one parallel processing unit that executes atleast one of Single Instruction Multiple Data instructions and Multipleinstruction Multiple Data instructions.
 24. The method of claim 14wherein the Doppler processing further comprises displaying at least oneof the presence, direction, and relative velocity of blood flow,perfusion and contour of a lumen within a body.
 25. The method of claim14 further comprising providing an adjustable control for at least oneof scan area selection, velocity display, steering angles, colorinversion, color gain, color priority, color persistence, color baselineand frame rate.
 26. The method of claim 14 further comprisingdetermining a phase shift value and optionally setting the phase shiftto zero.
 27. The method of claim 26 further comprising applying aspatial low pass filter to smooth phase shift values.
 28. The method ofclaim 27 further comprising converting phase shift data to color indexdata.
 29. The method of claim 26 wherein the phase shift is set to zeroif the power of a Doppler signal prior to application of a high passfilter is larger than a threshold value.
 30. The method of claim 14further comprising adjusting color priority with an adjustable control.31. A portable ultrasound imaging apparatus comprising: a batterypowered, handheld display device including a touchscreen display and adata processing module that is configured to perform a down conversionoperation, a scan conversion operation and a Doppler processingoperation, the display device being connected to a transducer array; amemory operably coupled to the data processing module, wherein thememory stores executable instructions that cause the data processingmodule to: select a pulse repetition frequency using a first controlwindow of a graphical user interface that is displayed on thetouchscreen display; calculate a power level along a sample line withthe data processing module and compare the power level to a threshold;apply a programmable filter to ultrasound image data including a wallfilter value selected from a plurality of wall filter values displayedwith a second control window of the graphical user interface, theselected wall filter value being a function of the pulse repetition rateto generate filtered Doppler image data; calculate, in response toexecuting instructions with the data processing module, a plurality ofautocorrelation values using the filtered Doppler image data; calculate,in response to executing instructions with the data processing module, aphase shift of an autocorrelation phase to generate a monotonic functionfor all values of the phase shift corresponding to a range of Dopplervelocities; convert, in response to executing instructions, thecalculated phase shift to an index with the data processing module; anddisplay a plurality of images including at least one of color Doppler,directional power Doppler, power Doppler and pulsed wave Doppler on thetouchscreen display of the handheld display device.
 32. The apparatus ofclaim 31 wherein the apparatus further comprises a handheld probeoperable to transmit an ultrasound image and to obtain data indicativeof the flow characteristics of the region of interest.
 33. The apparatusof claim 31 further comprising adjustable controls for one of at leastscan area selection, velocity display, steering angles, color inversion,color gain, color priority, color persistence, color baseline and framerate.
 34. The apparatus of claim 31 wherein the programmable filtercomprises a high pass filter.
 35. The apparatus of claim 31 furthercomprising a transducer array connected to the handheld display devicethat comprises a laptop or palmtop device having a plurality of parallelprocessing units.
 36. The apparatus of claim 35 wherein the handhelddisplay device includes a beamformer and a system controller.
 37. Thesystem of claim 1 wherein the handheld processor housing comprises alaptop computer.
 38. The system of claim 1 wherein the handheldprocessor housing comprises a palm held computer.
 39. The method ofclaim 14 wherein the graphical user interface is operable to select aDoppler scan area within a displayed image wherein an insonating beamemitted by a transducer array is at an angle of 60 degrees or lessrelative to blood flow within the Doppler scan area.
 40. The method ofclaim 39 further comprising selecting a frame rate in response to aselected scan area.